Nuclear Medicine imaging techniques enable to acquire functional information on a patient's specific organ or body system. This functional information is attained from analysis of internal radiation from pharmaceutical substance administered to the patient, which is labelled with a radioactive isotope. The radioactive isotope decays, resulting in the emission of gamma rays, thus providing information on the concentration of the radiopharmaceutical substance in regions of the patient's body. An instrument for the detection of gamma ray emissions of the radiopharmaceutical substance administered in the body is known as gamma camera. The Gamma camera collects gamma ray photons that are emitted from the patient's body, and the collected data is used to reconstruct an image or a series of images of the place in the body from which the gamma rays are originated. From this picture a physician can determine how a particular organ or system is functioning.
The main components making up a conventional gamma camera are photon detector crystal or detector array, a collimator for limiting the detection of incident gamma rays to a predetermined view angle, position logic circuits and data analysis computer. Depending on the type of the detector crystal, conventional gamma camera may or may not include a photo-multiplier tube array.
A gamma ray photon that has passed through the collimator interacts with the detector crystal by means of the Photoelectric Effect or Compton Scattering with ions of the crystal. These interactions cause the release of electrons, which in turn interact with the crystal lattice to produce light, in a process known as scintillation. Since only a very small amount of light is given off from the crystal, photo-multiplier tubes are normally attached to the back of the crystal. Typically, a conventional gamma camera has several photo-multiplier tubes arranged in a geometrical array. The position logic circuits that follow the photo-multiplier tube array receive electrical impulses from the tubes and determine where each scintillation event occurred in the detector crystal. Finally, in order to deal with the incoming projection data and to process it into a meaningful image of the spatial distribution of activity within the patient, a processing computer is used. The computer may employ various processing methods to reconstruct an image.
Different collimators are used in gamma cameras to limit the detection of photons to incidence range of predetermined angles. A parallel-hole collimator is typically made of lead or tungsten and has thousands of straight parallel holes in it, allowing only those gamma rays travelling in certain directions to reach the detector. As a result, the ratio of emitted versus detected photons may reach as high as 10000 to 1. In order to decrease this ratio, converging or diverging hole collimators, for example, fan-beam and cone-beam are also known in the art. The usage of these collimators increases the number of photon counts, which consequently improves sensitivity. Sensitivity, however, is inversely related to geometric resolution, which means that improving collimator resolution (i.e., having smaller diameter holes) decreases collimator sensitivity, and vice versa.
Single photon emission (SPE) imaging is a known nuclear medicine imaging technique. Several modes of SPE imaging are in use:
One of them is Single Photon Emission Computerized Tomography (SPECT). In this technique the gamma camera is rotated around the region of interest in the patient's body, and data is collected at several angular positions (hereafter referred to as angular projections). A fully three dimensional image is reconstructed from these angular positions.
SPECT is considered to be a very useful technique and a good tool for obtaining functional diagnostic information, however it requires the collection of large number of emitted photons (large statistics) and this means that in order to obtain the required number of photons, a long acquisition time is necessary. Long acquisition time means that the patient is subjected to a relatively long period of discomfort, and, furthermore, the overall number of patients who can be imaged in a given time is relatively small—a feature that many medical institutes and hospitals regard as an extremely unfavourable and undesirable situation.
Multi-detector cameras are capable of acquiring more photons per second than single detector cameras. Most popular are dual-detector cameras in which the two detectors are mounted in two different positions around the patient and are rotated simultaneously around him. The effective acquisition time of a dual-detector camera is that of a single-detector camera. For example, acquisition time of 10 minutes by a dual detector camera will result in effective acquisition time of 20 minutes.
Filtered Back Projection (FBP) is the most popular method of reconstruction of three-dimensional image from the acquired data set of angular projections. This method requires relatively short calculation time and is readily available commercially. FBP algorithms suffer from image quality degradation when the number of angular projections is low or when the angular distribution is irregular.
Nuclear emission is a stochastic process, thus the relative statistical noise associated with each acquired angular projection increases as the number of acquired photons used to form the projection decreases. Applying FBP algorithm on data set that includes high statistical noise causes degradation to the quality of the reconstructed image.
The current use of collimators results in a rather low detection efficiency of conventional SPECT, which leads to a prolonged data acquisition time and the need to administer high dosage of the radiopharmaceutical substance. The dosage used is determined by maximal radiation that can be safely tolerated by a patient.
Generally, each detector of a gamma camera is capable of being equipped with one of several interchangeable collimators, which are changed in order to match the type of diagnostic procedure. The collimators used in standard size gamma camera are heavy and costly.
The combination of dosage, collimator type and acquisition time, which sets image quality, that is adequate for specific diagnostic procedure was determined by nuclear imaging professionals during years of experience.
Another parameter that strongly affects both data acquisition time and image quality is the number of pixels in the image.
In pixilated detectors, such as solid-state detectors, the number of pixels in the reconstructed image is limited by the number of pixels of the detector. In scintillation-based detectors, the acquired data undergoes discretization process in which the photons are binned according to their position on the detectors. Similarly, the reconstruction of a three dimensional image is also set to a finite matrix of voxels. Since the reconstruction process often includes Fourier Transformation, the matrix size is generally chosen to be in the form of 2m where m is an integer. Practical matrix sizes that are typically used are 64×64 and 128×128.
The discretization process may be done during the data acquisition. Alternatively, the location on the detector of each impinging photon may be saved at full spatial resolution in a file as a list of events, optionally with additional information such as its energy or the position of the detector or the time of the event. Later, the list of events is analyzed and the discretization process is performed off line. This method of data collection and processing is known as List-Mode acquisition.
Due to the discretization process, the average number of acquired photons per pixel is inversely proportional to the number of pixels in the matrix:Av=Num/(2m)2
Where Num is the number of photons acquired and Av is the average number of photon in a pixel.
Since radioactive emission is a stochastic process, the relative stochastic noise associated with the number of photons counted in each pixel decreases as the number of counted photon increases. Image quality is strongly adversely affected by that noise. Thus, in situations where the number of acquired photons is small, such as in short duration acquisition, a smaller matrix size is chosen.
However, the resolution of the image is limited by the size of the pixels, so that larger matrices are desirable in situations where small features in the image have to be resolved.
Iterative reconstruction methods are used for SPECT. PCT/IL01/00730, published as WO 02/12918, presently allowed U.S. application Ser. No. 10/333,947, filed Jan. 22, 2003, and incorporated herein by reference, discloses methods for image reconstruction that result in enhanced three-dimensional nuclear image. These methods make uses and take advantage of collimators whose sensitivity is higher than collimators traditionally used in hospitals in order to collect larger number of detected photons within the data set, used for reconstruction of the enhanced three-dimensional image of superior quality.
In cardiac SPECT imaging the radiopharmaceutical distribution in the myocardium of a patient is imaged. Since the heart is beating, the heart wall motion blurs the image that is reconstructed from the accumulated data. In Gated SPECT imaging, the imaging is synchronized with the heart movement-cycle using electrocardiogram (ECG) signal.
In gated SPECT, each angular projection is divided to S sub-projections. All the sub-projections belonging to the same projection are acquired at the same angle, but at different times. Using the patient's ECG signal, the heartbeats are detected and the time between the end diastolic and end systolic phases is divided by S segments (typically S=8). Data acquired in each time segment is accumulated in the corresponding sub-projection.
Image reconstruction is done separately for the S groups of sub-projections, each showing different and relatively un-blurred phase of the heart motion. In addition, it is customary to reconstruct an image of the totality of all the acquired data. This image, although blurred by the motion, has less statistical noise. U.S. Pat. No. 6,507,752, incorporated herein by reference, for example, demonstrates a method of quantitative determination of cardiac muscle control by electrocardiogram synchronized traverse tomogram.
Tissue of various organs in the patient body attenuate gamma photons. This attenuation causes degradation of image quality due to loss of photons and more importantly, image distortion due to varying amount of attenuation between different organs of the body and the detector as it views the body from various direction.
Attenuation correction methods may be used to correct the distortion caused by tissue gamma absorption. However, in order to address the attenuation problem, an attenuation map—a quantities description of the attenuation within the patient's body is needed. One method for obtaining patient attenuation map is to position the patient between a radioactive source and the detector so that gamma photons pass through the patient body and get detected by the detector. The patient body is then scanned and angular attenuation projections are calculated from comparing the number of photons transmitted through the patient to the number of photons detected in the absence of the patient. Since attenuation through the dense parts of the body is considerably high the number of transmitted photons is low and thus image quality of the reconstructed attenuation map is limited by the time spent on acquiring the transmission data.
When emission data is gated, attenuation data may have to be gated as well to reflect the motion of the patient chest during heartbeats. For example, U.S. Pat. No. 6,429,434, incorporated herein by reference teaches a technique to address the attenuation problem.